Research on neural-machine interfaces (NMI) in recent years has demonstrated the feasibility of driving motor prosthesis for the upper limbs. Neural interface reliability has been identified as a critical research area where progress is needed prior to transitioning NMI technology for practical restoration of motor function in humans. Two key issues are 1) the inability of current interfaces to reliably obtain accurate information from neurons over a period of decades, and 2) currently measured neural signals cannot be reliably used to control prostheses with high speed and resolution.
Neural probe hardware implanted in the brain tissue is a critical element in achieving these reliability goals. Failure of neural probes may be caused by several issues. After implantation, current probes are surrounded by reactive microglia and reactive astrocyte scarring as shown pictorially in FIG. 1A. Tissue reaction with the probe results in encapsulation that insulates the electrode by impeding diffusion and may impede current flow. Encapsulation increases the distance of the electrode from active neurons. For viable recording, the distance of the electrode from active neurons must be less than 100 μm. Progressive death and degeneration of neurons in the zone around the inserted probe due to chronic inflammation may eliminate neural electrophysiological activity. Lastly, interconnects may fatigue and break due to stresses. Experiments in animals have resulted in some neural electrode sites failing while others keep working for several years. This variability in outcome is believed to be due to several factors including variable blood-brain barrier (BBB) damage, variable scar formation, mechanical strain from micromotion, inflammation, microglial condition and disconnected neurons.
Neural probes employed today for neural-machine interface studies are essentially stiff needles usually made from wires, silicon or glass. Metal wire neural probes are typically 50-100 μm in diameter and usually made of platinum or iridium and insulated with glass, Teflon, polyimide or parylene. Basic tests were performed in 1974 determining that iridium, rhodium, platinum and palladium, in that order, have excellent resistance to electrolysis under conditions simulating biostimulation applications Iridium appears best and microelectrodes should exhibit lifetimes against electrolysis of decades. 2D arrays of wire probes have been made for chronic implantation. Drawbacks of this approach are the manual assembly, the lack of multiple electrodes per shaft, and issues with the predominantly metal wire splaying when inserted in the tissue. FIG. 1B is a predicted outcome with the miniature, ultra-compliant probe according to the present invention.
Silicon probes made with MEMS fabrication were first introduced by Ken Wise and Jim Angell at Stanford in 1969. Ken Wise's group at the University of Michigan subsequently developed a series of silicon probes and probe arrays with multi-site electrodes. The Michigan probes are made through a wet etch step that stops on boron-doped Si and necks the shank thickness down to around 15 μm. In more recent work, Si Deep reactive-ion etching (DRIE) has been used to make Si probes without the boron etch stop.
A 2D probe array was developed at the University of Utah in 1991, known as the Utah Electrode Array (UEA). The shanks in the UEA are made by sawing grooves into the substrate followed by a silicon wet etch to smooth the sidewalls and sharpen the needles. Platinum is deposited on the needles, which are subsequently coated in polyimide with just the tip exposed. Iridium has also been used for metallization. The UEA has been demonstrated repeatedly to record chronically, has recording sites 50 to 100 μm long, suggesting to some researchers that large recording site sizes perform better for chronic recording.
Polycrystalline diamond (poly-C) probes with 3 μm thick undoped poly-C on a ˜1 μm SiO2 layer have been fabricated by Dr. Aslam's group at Michigan State University. These probes, 5 mm long, are capable of electrical and electrochemical recording with AgCl reference electrodes, Au counter electrodes and doped poly-C working electrodes.
Research groups have created more compliant probes made with thin-film wiring embedded in polymer insulating films. Flexible central nervous system (CNS) probes have been made in polyimide, SU8/parylene and all parylene. However, flexible probes for single-unit detection have had to maintain adequate stiffness for insertion into the brain tissue. Thus, all prior art probes are made with a straight shank and relatively large probe diameters. The result is that even the most advanced of today's probes are extremely stiff in both axial and transverse directions relative to brain tissue, which has a Young's modulus of approximately 30 kPa. Any axial force transmitted through the external cabling directly acts on the probe and creates shear forces at the electrode-tissue interfaces. Such forces may come from external motion or from tissue growth around the implant. A team from Drexel Univ., the Univ. of Kentucky and SUNY created ceramic-based multisite microelectrode arrays on alumina substrates with thickness ranging from 38 to 50 μm, platinum recording sites of 22 μm×80 μm, and insulation using 0.1 μm ion-beam assisted deposition of alumina. The 7 mm-long shanks widen to 700 μm at the base. Their experience has shown that if the electrode is implanted slowly, there is a greater likelihood of recording a single neuron and when implanted quickly single units could not be recorded. After 3 months, there were no clear single neuron recordings from any recording sites and subsequent immunohistology showed glial formation for several hundred microns beyond the insertion hole.
Y.-C. Tai's group at Caltech produced parylene-coated silicon probes with integral parylene cabling, shown in FIG. 2A. The shanks were up to 12 mm long. Parylene adhesion to silicon was enhanced by performing a short XeF2 etch to roughen the substrate. A primary innovation was a flexible 10 μm-thick, 830 μm-wide, 2.5 mm-long parylene cable. Sacrificial photoresist was placed under the cable as a release layer. Backside Si DRIE was used to free the device from the rest of the substrate. A 200° C. anneal for 48 hr was performed to soften and allow straightening of the meanders in the cabling from the compact planar layout. 16-channel Omnetics connectors were bonded with conductive epoxy in through-holes on the parylene and backed with a printed circuit board for mechanical strength.
Flexible polyimide probe arrays (FIG. 2C) have been made with gold electrodes. These probes must be inserted by first creating an insertion hole with a scalpel or needle. A later polyimide probe array incorporated silicon for selected locations along the length of the shank, with polyimide connectors to create enhanced compliance, as shown in FIG. 2B.
FIGS. 3A and 3B illustrated an all-polymer probe design incorporated a lateral lattice-like parylene structure attached to a larger SU8 shank to reduce the structural size close to the electrodes. The lattice structure, shown in FIG. 3A, included a 4 μm-wide, 5 μm-thick lateral beam (See FIG. 3D) located parallel to the main shank. Encapsulating cell density around the lateral beam was reduced by one-third relative to the larger shank. SU8 shank can be coated with a parylene coating forming a coated shank having a depth of 48 μm.
Flexible probes have been made in all parylene with 0.5 μm-thick gold interconnect and electrodes. The parylene cross-sectional dimension was set at 100 μm wide and 25 μm thick to create adequate stiffness for insertion. Probes up to 2.5 mm long, shown in FIG. 2D, were designed but insertion was performed to 0.5 mm. The probe retained mechanical and electrical integrity following acute tests in rats after controlled cortical impact with a 2 mm stroke. Researchers at IMTEK created ultra compliant ECoG probes on a polyimide/platinum foil substrate. The 252-channel array spanning 35 mm by 60 mm was made by spinning two 5 mm-thick polyimide films onto a silicon substrate. The first polyimide film was roughened by O2 plasma to enhance adhesion of the direct sputtered 300 nm-thick platinum interconnect layer. Tweezers were used to pull the finished array from the wafer. Omnetics connectors were electrically bonded to through-holes in the film with solder paste and mechanically fixed with epoxy. QuinetiQ is developing liquid crystal polymer neural probes and cables. Rogers' group at UIUC has developed implantable electrodes using transfer of silicon parts onto bioresorbable silk with a Polydimethylsiloxane (PDMS) stamp process. This approach will prove useful for ECoG electrodes as well.
U.S. patent application 20090099441 from Dr. Giszter's Drexel group describes biodegradable stiffening wires 1 braided with electrode wires 2 with electrodes 3 (see FIG. 3C) where flexible wires 2 are braided onto a large diameter, stiff maypole structure 4 with stiff biodegradable strands 1. When the biodegradable strands 1 dissolve, the flexible wiring 2 is left in the brain tissue. Reliable and manufacturable connections to the braided wires become difficult when scaled to arrays.
Olbricht et al has reported on flexible microfluidic devices supported by biodegradable insertion scaffolds for convection-enhanced neural drug delivery. The device consists of a flexible parylene-C microfluidic channel that is supported during its insertion into tissue by a biodegradable poly(DL-lactide-co-glycolide) (PLGA) scaffold. The scaffold is made separately by hot embossing the PLGA material into a mold. The parylene-C microfluidic channel is then manually assembled by first tacking it down to the scaffold with a drop of epoxy followed by a dip in dichloromethane to partly dissolve some of the PLGA and thereby attach it to the parylene-C shank. The PLGA shanks were nearly 100% degraded after 27 days in organic chemical buffering agent and were compliant after 15 to 18 days.
Suzuki, Mabuchi et al, describe multichannel flexible neural probes coated with PLGA microspheres that were infused with nerve growth factor. Two types of neural probes were created. Both probe types included flexible thin-film parylene-C probes. The first probe (type-A) included a parylene-C groove along the shank for manual placement of PLGA microspheres mixed with polyethylene glycol. The second probe (type-B) included multiple electrodes without the groove structure. The PLGA was manually coated to create the biodegradable shank for insertion. Both probes were inserted in a rat cortex, with successful neural recording from the type-A probe. Neural signals were not observed from the type-B probe, presumably due to residual PLGA obstructing the electrode.
Tyler et al, have developed a neural probe made from a polymer nanocomposite of poly(vinyl acetate) (PVAc) and tunicate whiskers, inspired by the sea cucumber dermis. The probe material exhibits a real part of the elastic modulus (tensile storage modulus) of 5 GPa after fabrication. When exposed to physiological fluid conditions, its modulus decreases to 12 MPa. The probe did not include wiring. Results from animal implantation studies showed an increased neuronal density and decreased glial formation around the PVAc probe when compared to a 50 μm-diameter tungsten wire probe. This work provides evidence that mechanical flexibility of the probe is an important aspect of reliable neural probes.